There have been broadly employed radiographic images such as X-ray images for diagnosis of the conditions of patients on the wards. Specifically, radiographic images using a intensifying-screen/film system have achieved enhancement of speed and image quality over its long history and are still used on the scene of medical treatment as an imaging system having high reliability and superior cost performance in combination. However, these image data are so-called analog image data, in which free image processing or instantaneous image transfer cannot be realized.
Recently, there appeared digital system radiographic image detection apparatuses, as typified by a computed radiography (also denoted simply as CR) and a flat panel radiation detector (also denoted simply as FPD). In these apparatuses, digital radiographic images are obtained directly and can be displayed on an image display apparatus such as a cathode tube or liquid crystal panels, which renders it unnecessary to form images on photographic film. Accordingly, digital system radiographic image detection apparatuses have resulted in reduced necessities of image formation by a silver salt photographic system and leading to drastic improvement in convenience for diagnosis in hospitals or medical clinics.
The computed radiography (CR) as one of the digital technologies for radiographic imaging has been accepted mainly at medical sites. However, image sharpness is insufficient and spatial resolution is also insufficient, which have not yet reached the image quality level of the conventional screen/film system. Further, there appeared, as a digital X-ray imaging technology, an X-ray flat panel detector (FPD) using a thin film transistor (TFT), as described in, for example, the article “Amorphous Semiconductor Usher in Digital X-ray Imaging” described in Physics Today, November, 1997, page 24 and also in the article “Development of a High Resolution, Active Matrix, Flat-Panel Imager with Enhanced Fill Factor” described in SPIE, vol. 32, page 2 (1997).
To convert radiation to visible light is employed a scintillator panel made of an X-ray phosphor which is emissive for radiation. The use of a scintillator panel exhibiting enhanced emission efficiency is necessary for enhancement of the SN ratio in radiography at a relatively low dose.
Generally, the emission efficiency of a scintillator panel depends of the scintillator thickness and X-ray absorbance of the phosphor. A thicker phosphor layer causes more scattering of emission within the phosphor layer, leading to deteriorated sharpness. Accordingly, necessary sharpness for desired image quality level necessarily determines the layer thickness.
Specifically, cesium iodide (CsI) exhibits a relatively high conversion rate of from X-rays to visible light. Further, a columnar crystal structure of the phosphor can readily be formed through vapor deposition and its light guide effect inhibits scattering of emitted light within the crystal, enabling an increase of the phosphor layer thickness.
However, the use of CsI alone results in reduced emission efficiency. For example, Japanese Patent Publication JP-B 54-35060 disclosed a technique for use as an X-ray phosphor in which a mixture of CsI and sodium iodide (NaI) at any mixing ratio was deposited on a substrate to form sodium-activated cesium iodide (CsI:Na); and recently, a mixture of CsI and thallium iodide (TlI) at any mixing ratio was deposited on a substrate to form thallium-activated cesium iodide (CsI:Tl), which was further subjected to annealing as a post-treatment to achieve enhanced visible-conversion efficiency.
There were also proposed other means for enhancing light output, including, for example, a technique of rendering a substrate to form a scintillator thereon reflective; a technique of forming a reflection layer on a substrate; and a technique of a scintillator on a transparent organic film covering a reflective thin metal film provided on a substrate. There was also known a technique in which an activator concentration was made uniform at any portion of a phosphor layer, resulting in enhanced emission luminance (as described in, for example, Patent document 1.
However, there has been desired a scintillator panel of enhanced emission efficiency.
There were also known a technique in which a scintillator was coated on columnar crystals formed on a FOP substrate to emit two color luminescence (as described in, for example, Patent document 2) and a technique of using a scintillator panel provided with a phosphor layer having a plural-layered structure (as described in, for example, Patent document 3), either of which was not aimed at enhancement of emission luminance.
Patent document 1: JP 2003-28994A
Patent document 2: JP 2005-106541A
Patent document 3: JP 2005-148060A